Proportional pressure assist ventilation controlled by a diaphragm electromyographic signal

ABSTRACT

A closed loop system uses (a) the intensity of the diaphragm electromyogram (EMG) for a given inspiratory volume; (b) the inspiratory volume for a given EMG intensity; or (c) a combination of (a) and (b); in view of controlling the level of gas flow, gas volume or gas pressure delivered by a mechanical (lung) ventilator. The closed loop ventilator system enables for automatic or manual adjustment of the level of inspiratory support in proportion to changes in the neuro-ventilatory efficiency such that the neural drive remains stable at a desired target level. An alarm can also be used to detect changes in neuroventilatory efficiency in view of performing manual adjustments.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to a system using the intensity of thediaphragm electromyogram (EMG) at a given lung volume or the lung volumeat a given EMG intensity to automatically or manually adjust the levelof inspiratory support in proportion to changes in the neuro-ventilatoryefficiency.

The present invention also relates to a system responsive to theintensity of the diaphragm electromyogram (EMG) measured immediatelybefore the onset of inspiratory flow to automatically or manuallycontrol and maintain an optimum level of extrinsic positive endexpiratory pressure (PEEP) applied to a patient, and to automatically ormanually control a duration from the onset of EMG to onset ofrespiratory flow.

2. Brief Description of the Prior Art

Prior art algorithms used to create closed-loop ventilator systems arebased on variables such as tidal volume, respiratory rate, inspiratoryflow, end-tidal carbon dioxide levels and/or rate of rise in pressure.However, none of these parameters can provide a reliable measure of therespiratory neural drive because they are affected by changes inneuro-mechanical or neuro-ventilatory efficiency.

Neuro-ventilatory efficiency is a term used to express the amount ofneural drive (breathing effort) needed to obtain a given tidal lungvolume. In brief, neural drive is converted into mechanical tension, aprocess which is influenced by the muscle length, temperature,electrolyte imbalance, etc. The role of inspiratory flow in the linkbetween neural drive and mechanical tension has previously beensuggested; however the proposed influence could not be demonstrated formean inspiratory flow rates up to 1.4 liters/second. The mechanicaltension is then translated into pressure. a process which is affected bythe shape of the diaphragm dome. Finally the pressure expands thealveoli and causes air to flow, and the translation of pressure tovolume depends on the elasto-viscous behaviour of the respiratorysystem. Consequently, there are many factors that may influence thetidal volume output obtained for a given increase in neural drive(inspiratory effort).

Evaluation of respiratory drive by measurements such as the rate of risein pressure or lung volume is not reliable when, for example, the musclelength or the respiratory system impedance are affected by changes inthe neuro-ventilatory efficiency. In a patient, airway resistance andelastance can change from one minute to another and muscle length iscontinuously altered.

OBJECTS AND SUMMARY OF THE INVENTION

An object of the present invention is therefore to eliminate thedrawbacks of the prior art.

Another object of the present invention is to provide a closed loopsystem using:

-   (a) the intensity of the diaphragm electromyogram (EMG) for a given    inspiratory volume;-   (b) the inspiratory volume for a given EMG intensity; or-   (c) a combination of (a) and (b);    in view of controlling the level of gas flow, gas volume or gas    pressure delivered by a mechanical (lung) ventilator; the closed    loop ventilator system enables for automatic or manual adjustment of    the level of inspiratory support in proportion to changes in the    neuro-ventilatory efficiency such that the neural drive remains    stable at a desired target level. An alarm can also be used to    detect changes in neuroventilatory efficiency in view of performing    manual adjustments.

Another object of the present invention is to provide a closed-loopsystem responsive to the intensity of the diaphragm EMG measuredimmediately before the onset of inspiratory flow to quantifypre-inspiratory breathing effort in view of automatically or manuallyadjusting a level of extrinsic positive end expiratory pressure (PEEP)applied to a patient in proportion to changes in EMG intensity ofpre-inspiratory efforts. In this manner, the pre-ventilatory intensityof the diaphragm EMG can be maintained at a desired, minimum level suchthat the pre-inspiratory neural drive remains stable at a desired targetminimal level. Determination of the duration from the onset of EMG tothe onset of respiratory flow is also used for quantitative evaluationof the intrinsic PEEP, and to guide adjustment of the triggersensitivity of the ventilator systems.

Different from pressure and ventilatory related indexes, the intensityof the EMG represents the temporal (mean MU (motor unit) rate coding)and spatial (MU recruitment) summation of action potentials and isobtained at the level of the sarcolemma muscle. The intensity of the EMGis therefore not affected by changes in the muscle's neuro-ventilatorycoupling. In the present invention, the use of crural diaphragm EMGrests on the assumption that neural drive to the crural diaphragm isrepresentative for the total respiratory drive. It is also based on thecondition that neuromuscular transmission and innervation of the cruraldiaphragm are normal. For breathing with increased demand thisassumption is well founded. Hence, the intensity of the EMG needed toproduce a given inspiratory volume should express the efficiencyrelation between neural drive and volume output.

The objects, advantages and other features of the present invention willbecome more apparent upon reading of the following non restrictivedescription of a preferred embodiment thereof, given by way of exampleonly with reference to the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

In the appended drawings:

FIG. 1 is a schematic representation of a set-up of an EMG analysissystem;

FIG. 2 is a section of oesophageal catheter on which an array ofelectrodes of the EMG analysis system of FIG. 1 is mounted;

FIG. 3 illustrates a section of oesophageal catheter on which a secondembodiment of the array of electrodes is mounted;

FIG. 4 is a graph showing a set of EMG signals of the diaphragm (EMGdisignals) detected by pairs of successive electrodes of the array of FIG.2;

FIG. 5 is a flow chart showing a method for conducting a doublesubtraction technique of the EMGdi signals;

FIG. 6 is a graph showing the distribution of correlation coefficientscalculated for determining the position of the center of thedepolarizing region of the diaphragm along the array of electrodes ofFIG. 2;

FIG. 7 is a schematic diagram illustrating in the time domain a doublesubtraction technique for improving the signal-to-noise ratio and toreduce an electrode-position-induced filter effect along the array ofelectrodes of FIG. 2;

FIG. 8 a is a graph showing the power density spectrum of electrodemotion artifacts, the power density spectrum of ECG, and the powerdensity spectrum of EMGdi signals;

FIG. 8 b is a graph showing an example of transfer function for a filterto be used for filtering out the electrode motion artifacts, ECG, andthe 50 or 60 Hz disturbance from electrical mains;

FIG. 9 is a schematic diagram illustrating in the frequency domainstabilization by the double subtraction technique of the centerfrequency upon displacement of the center of the depolarizing region ofthe diaphragm along the array of electrodes of FIG. 2;

FIG. 10 is a schematic block diagram of a system according to theinvention for controlling inspiratory assist by means of an EMGdi signalobtained with the above mentioned double subtraction technique and ameasurement of the volume of air breathed by the patient by acommercially available system;

FIG. 11 is a schematic block diagram of a system according to theinvention (a) capable to determine the time delay from the onset of EMGto the onset of inspiratory flow and (b) using the level ofpre-inspiratory effort obtained through the EMGdi signal intensity(common noise level subtracted) during a predetermined time periodimmediately preceding the onset of inspiratory flow to indicate thepresence of “intrinsic PEEP” and to adjust the level of applied“extrinsic PEEP” and/or ventilator trigger sensitivity such that thelevel of pre-inspiratory effort is suppressed, i.e the EMGdi signalintensity (common noise level subtracted) during the above mentionedpredetermined time period is close to zero;

FIG. 12 a is an exemplary graph of a patient's inspiratory flow versustime for quiet breathing in COPD (Chronic Obstructive PulmonaryDisease); and

FIG. 12 b is an exemplary graph of a patient's EMG RMS intensity versustime for quiet breathing in COPD.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

Although the preferred embodiment of the present invention will bedescribed in relation to a double subtracted EMGdi signal, it should bekept in mind that the concept of the present invention can be used withany respiratory muscle signal.

To measure EMG activity of the diaphragm 11 (EMGdi) of a human patient14, an array of electrodes such as 12 (FIGS. 1 and 2) are mounted on thefree end section 15 of an oesophageal catheter 13, with a constantinter-electrode distance d (FIG. 2). As shown in FIG. 1, the catheter 13is introduced into the patient's oesophagus through one nostril or themouth until the array of electrodes 12 is situated at the level of thegastroesophageal junction. The diaphragm 11 and/or the oesophagusslightly move during breathing of the patient 14 whereby the array ofelectrodes 12 also slightly moves about the diaphragm 11. As will beexplained in the following description, autorhatic compensation for thisdisplacement is provided for.

According to a preferred embodiment, an electrode 12 is mounted on thefree end section 15 of the catheter 13 by winding stainless steel wire(not shown) around that catheter 13. The wound stainless steel wirepresents a rough surface smoothed out by solder, which in turn iselectroplated with nickel, copper and then gold or silver. Of course, itis within the scope of the present invention to use other electrodestructures. Also, the electrodes 12 can possibly be applied to anasogastric feeding tube (not shown) which is routinely introduced inintensive-care unit (ICU) patients.

Electric wires (not shown) interconnect each pair of successiveelectrodes such as 1–7 (FIG. 2) with a respective one of a group ofdifferential amplifiers 16. Obviously, these electric wires follow thecatheter 13 from the respective electrodes 12 to the correspondingamplifiers 16, and are preferably integrated to the catheter 13.Preferably, the electric wires transmitting the EMGdi signals collectedby the various pairs 1–7 of electrodes 12 are shielded to reduce theinfluence of external noise, in particular disturbance from the 50 or 60Hz current and voltage of the electrical mains.

The group of differential amplifiers 16 amplifies (first subtractionstep of a so-called double subtraction technique) and band-pass filterseach EMGdi signal. This first subtraction step may also be carried outin the personnal computer 19 when the amplifiers 16 are single-ended orequivalently designed amplifiers (monopolar readings).

In the example illustrated in FIGS. 1 and 2, the free end section 15 ofthe catheter 13 is provided with an array of eight electrodes 12defining seven pairs 1, 2, 3, 4, 5, 6 and 7 of successive electrodes 12respectively collecting seven different EMGdi signals. Although it hasbeen found that EMG activity of the diaphragm (EMGdi) can be measuredaccurately with an oesophageal catheter 13 provided on the free endsection 15 thereof with an array of eight electrodes 12, a differentnumber and/or configuration of pairs of electrodes 12 can becontemplated depending on the patient's anatomy and movement of thediaphragm. Also, the pairs 1–7 do not need to be pairs of successiveelectrodes; as an example FIG. 3 illustrates an array of nine electrodesto form seven overlapping pairs of electrodes 1–7.

A major problem in recording EMGdi signals is to maintain the noiselevel as low and as constant as possible. Signals the electric wirestransmitting the EMGdi signals from the electrodes 12 to thedifferential amplifiers 16 act as an antenna, it is crucial, asindicated in the foregoing description, to shield these electric wiresto thereby protect the EMGdi signals from additional artifactual noise.Also, the package enclosing the differential amplifiers 16 is preferablymade as small as possible (miniaturized) and is positioned in closeproximity to the patient to decrease as much as possible the distancebetween the electrodes 12 and the amplifiers 16.

The amplified EMGdi signals are sampled by a personal computer 19through respective isolation amplifiers of a unit 18, to form signalsegments of fixed duration. Unit 18 supplies electric power to thevarious electronic components of the differential and isolationamplifiers while ensuring adequate isolation of the patient's body fromsuch power supply. The unit 18 also incorporates bandpass filtersincluded in the respective EMGdi signal channels to eliminate theeffects of aliasing. The successive EMGdi signal segments are thendigitally processed into the personal computer 19 afteranalog-to-digital conversion thereof. This analog-to-digital conversionis conveniently carried out by an analog-to-digital converterimplemented in the personal computer 19. The personal computer 19includes a monitor 40 and a keyboard 31.

It is believed to be within the capacity of those of ordinary skill inthe art to construct suitable differential amplifiers 16 and an adequateisolation amplifiers and power supply unit 18. Accordingly, theamplifiers 16 and the unit 18 will not be further described in thepresent specification.

An example of the seven EMGdi signals collected by the pairs 1–7 ofsuccessive electrodes 12 (FIGS. 1 and 2) and supplied to the computer 19is illustrated in FIG. 4.

As the diaphragm is generally perpendicular to the longitudinal axis ofthe oesophageal catheter 13 equipped with an array of electrodes 12,only a portion of the electrodes 12 are situated in the vicinity of thediaphragm. It is therefore important to determine the position of thediaphragm with respect to the oesophageal electrode array.

The portion of the crural diaphragm 11 which forms the muscular tunnelthrough which the oesophageal catheter 13 is passed is referred to the“diaphragm depolarizing region” (DDR). The thickness of the DDR is 20–30mm. It can be assumed that, within the DDR, the distribution of activemuscle fibers has a center from which the majority of the EMGdi signalsoriginate, i.e. the “diaphragm depolarizing region center” (DDR center).Therefore, EMGdi signals detected on opposite sides of the DDR centerwill be reversed in polarity with no phase shift; in other words, EMGdisignals obtained along the electrode array are reversing in polarity atthe DDR center.

Moving centrally from the boundaries of the DDR, EMGdi power spectrumsprogressively attenuate and enhance in frequency. Reversal of signalpolarity on either side of the electrode pair 4 with the most attenuatedpower spectrum confirms the position from which the EMGdi signalsoriginate, the DDR center.

Referring to FIG. 5, the first task of the computer 19 is to determinethe position of the center of the DDR along the array of electrodes 12.The center of the DDR is repeatedly determined at predetermined timeintervals.

For that purpose, filtering step 505 removes from each EMGdi signal themotion artifacts, the electrocardiogram (ECG) component, and thedisturbance from the electrical mains. Motion artifacts are induced bymotion of the electrodes 12. More generally motion artifacts are definedas a low frequency fluctuation of the EMGdi signals' DC level induced bymechanical alterations of the electrode metal to electrolyte interfacei.e. changes in electrode contact area and/or changes in pressure thatthe tissue exerts on the electrode.

In step 501, the filtered EMGdi signals from step 505 arecross-correlated in pairs. As well known to those of ordinary skill inthe art, cross-correlation is a statistical determination of the phaserelationship between two signals and essentially calculates thesimilarity between two signals in terms of a correlation coefficient r(step 502). A negative correlation coefficient r indicates that thecross-correlated signals are of opposite polarities.

FIG. 6 shows curves of the value of the correlation coefficient r versusthe midpoint between the pairs of electrodes from which the correlatedEMGdi signals originate. In this example, the inter-electrode distanceis 10 mm. Curves are drawn for distances between the correlated pairs ofelectrodes 12 of 5 mm (curve 20), 10 mm (curve 21), 15 mm (curve 22) and20 mm (curve 23). One can appreciate from FIG. 5 that negativecorrelation coefficients r are obtained when EMGdi signals fromrespective electrode pairs situated on opposite sides of the electrodepair 4 are cross-correlated. It therefore appears that the change inpolarity occurs in the region of electrode pair 4, which is confirmed bythe curves of FIG. 4. Accordingly, it can be assumed that the center ofthe DDR is situated substantially midway between the electrodes 12forming pair 4.

For example, the center of the DDR can be precisely determined byinterpolation (step 503 of FIG. 5) using a square law based fit of thethree most negative correlation coefficients of curve 21 obtained bysuccessive cross-correlation of the EMGdi signal segments from eachelectrode pair to the EMGdi signal segments from the second nextelectrode pair. Association of the center of the DDR to a pair ofelectrodes 12 provides a “reference position” from which to obtain EMGdisignal segments within the DDR. Such control is essential in overcomingthe artifactual influence of perpendicular bipolar electrode filteringon the EMGdi power spectrum.

It has been experimentally demonstrated that EMGdi signals recorded inthe oesophagus are satisfactory as long as they are obtained fromelectrode pairs (with an inter-electrode distance situated between 5 and20 mm) positioned at a distance situated between 5 and 30 mm on theopposite sides of the DDR center (the inter-pair distance beingtherefore situated between 5 and 30 mm). Although EMGdi signals obtainedfrom these positions offers a clear improvement in acceptance rate, thesignal-to-noise ratio during quiet breathing still tends to remainunsatisfactorily low. The EMGdi signal obtained from one electrode pair(for example channel 0 in FIG. 7) situated in between the two electrodepairs used to produce the double subtracted signal, can be added to thisdouble subtracted signal either before as a raw signal or after when RMSor equivalent EMGdi signal measure has been computed, in order tominimize loss of signal.

For example, in FIG. 4, the EMGdi signals originating from the electrodepairs 3 and 5 situated respectively 10 mm below and 10 mm above the DDRare strongly inversely correlated at zero time delay. In contrast to theinversely correlated EMGdi signals, the noise components for electrodepairs 3 and 5 are likely to be positively correlated. Hence, asillustrated in FIG. 7, subtraction of the EMGdi signals 24 and 25 fromelectrode pairs 3 and 5 will result into an addition of thecorresponding EMGdi signals (signal 26 of FIG. 6) and into asubtraction, that is an elimination of the common noise components. Thistechnique will be referred to as “the double subtraction technique”(step 504 of FIG. 5). Again, the EMGdi signal obtained from oneelectrode pair (for example channel 0 in FIG. 7) situated in between thetwo electrode pairs used to produce the double subtracted signal, can beadded to this double subtracted signal either before as a raw signal orafter when RMS or equivalent EMGdi signal measure has been computed, inorder to minimize loss of signal.

Subtraction step 504 (second subtraction step of the double subtractiontechnique) can be carried out either in the time domain, or afterconversion of signals 24 and 25 in the frequency domain. Doublesubtraction technique can be performed by subtracting other combinationsof signals, for example by subtracting the EMGdi signal segments fromelectrode pair 2 from the EMGdi signal segments from electrode pair 5(FIG. 4), by subtracting signal segments from electrode pair 6 from thesignal segments from electrode pair 3 and by adding these differences,etc. What is important is to subtract two signals of opposite polaritiesobtained in the vicinity of the muscle. More than two signal pairs ofopposite polarities can be used in the double subtraction. Again, theEMGdi signal obtained from one electrode pair (for example channel 0 inFIG. 7) situated in between the two electrode pairs used to produce thedouble subtracted signal, can be added to this double subtracted signaleither before as a raw signal or after when RMS or equivalent EMGdisignal measure has been computed, in order to minimize loss of signal.

The double subtraction technique is carried out in step 504 on the pairof EMGdi signals (for example the signals from electrode pairs 3 and 6shown in FIG. 4) identified in step 503 after appropriate filtering ofthese EMGdi signals in step 506. Still again, the EMGdi signal obtainedfrom one electrode pair (for example channel 0 in FIG. 7) situated inbetween the two electrode pairs used to produce the double subtractedsignal, can be added to this double subtracted signal either before as araw signal or after when RMS or equivalent EMGdi signal measure has beencomputed, in order to minimize loss of signal.

The graph of FIG. 8 a shows the power density spectrum of the abovedefined electrode motion artifacts, the power density spectrum of ECG,and the power density spectrum of EMGdi signals. The graph of FIG. 8 bshows an example of transfer function for a filter (the dashed lineshowing the optimal transfer function, and the solid line the transferfunction implemented by the inventors) to be used in step 505 forfiltering out the electrode motion artifacts, ECG, and the 50 or 60 Hzdisturbance from the electrical mains. Processing of the EMGdi signalsby the computer 19 to follow as closely as possible the optimal transferfunction of FIG. 8 b will conduct adequately filtering step 505.

Therefore, double-subtracted signal segments 509 are obtained at theoutput of step 504 by subtracting the EMGdi signal segments from thepair of electrodes 12 in optimal location above the diaphragm from theEMGdi signal segments from the pair of electrodes 12 in optimal locationbelow the diaphragm. More than two signal pairs of opposite polaritiescan be used in the double subtraction. Again, the EMGdi signal obtainedfrom one electrode pair (for example channel 0 in FIG. 7) situated inbetween the two electrode pairs used to produce the double subtractedsignal, can be added to this double subtracted signal either before as araw signal or after when RMS or equivalent EMGdi signal measure has beencomputed, in order to minimize loss of signal.

Referring back to FIG. 5, step 506 calculates the RMS (root-mean-square)or equivalent or similar value 510 of the double-subtracted signalsegments 509 produced in step 504. The increase in intensity obtainedwith the double subtraction technique is associated with a twofoldincrease in RMS values. RMS values obtained with the double subtractiontechnique are closely and linearly related to the original signals. Itshould be kept in mind that the RMS value can be replaced by any othervalue representative of the strength of the double-subtracted signalsegments 509.

The digital RMS signal segment value 510 calculated by the computer 19in step 506 is finally digital-to-analog converted to an on-line analogRMS value 508 (step 507) in view of controlling a lung ventilator 54(FIG. 10). It should be mentioned that it is within the scope of thepresent invention to supply a digital value 608.

The double subtraction technique compensates for the changes in signalstrength and frequency caused by movement of the diaphragm 11 (FIG. 1)and/or the oesophagus during breathing of the patient 14 causingmovement of the array of electrodes 12 with respect to the diaphragm 11.Referring to FIG. 9, off center of the array of electrodes 12(electrode-position-induced filter effect) causes a variation of centerfrequency values due to filtering (see curves 27 and 28) for the EMGdisignals from the electrode pairs 3 and 5. The double subtractiontechnique eliminates such variation of center frequency values asindicated by curve 29 as well as variation of signal strength.Therefore, the reciprocal influence of the position of the DDR center onthe EMGdi signal frequency content is eliminated by the doublesubtraction technique.

It has been found that the double subtraction technique may improve thesignal-to-noise ratio by more than 2 dB and reduce anelectrode-position-induced filter effect. Double subtraction techniqueis also responsible for a relative increase in acceptance rate by morethan 30%.

Noise of non diaphragmatic origin or artifactual signals are stronglycorrelated at zero time delay and equal in polarity between all pairs ofelectrodes 12. Hence, this noise of non diaphragmatic origin orartifactual signals appear as a common mode signal for all electrodepairs and therefore, are substantially reduced by the double subtractiontechnique.

In the following description, it should be considered that the flow andvolume of air breathed by the patient can be measured by anycommercially available system.

Neuro-Ventilatory Efficiency;

The neuro-ventilatory efficiency is obtained by relating the diaphragmEMGdi signal intensity to changes in lung volume, or by relating thelung volume to changes in diaphragm EMGdi signal intensity. Since therelationship between the diaphragm EMGdi signal intensity and the lungvolume is not linear, this non-linearity is minimized by expressing:

-   the intensity of the diaphragm EMGdi signal for a given volume    change from end-expiratory lung volume, for example the EMGdi signal    intensity obtained during 400 ml inspiration starting from    end-expiratory lung volume (in the present disclosure, intensity is    intended to encompass the mean, peak, median and total RMS intensity    of the diaphragm EMGdi signal); or-   the lung volume obtained at a given diaphragm EMGdi signal    intensity.

A relatively small tidal lung volume is suitable because therelationship between diaphragm EMGdi signal intensity and lung volume isrelatively linear at this low range. Secondly, the use of a fixed, giventidal volume or diaphragm EMGdi signal intensity will protect againstthe non-linear influences and allows for a reliable estimation ofrelative changes in neuro-ventilatory efficiency.

In this manner, a ventilatory efficiency index expressing:

-   the EMGdi signal intensity for a given inspiratory lung volume    starting from the end-expiratory lung volume; or-   the lung volume for a given diaphragm EMGdi signal intensity;-   is calculated. If the EMGdi signal intensity for the above mentioned    given inspiratory lung volume or the lung volume for the above    mentioned given diaphragm EMGdi signal intensity is changing, the    above indicated index will also change and this change can be    expressed in percentage (%). For example, using the diaphragm EMGdi    signal intensity for the above mentioned fixed, given inspiratory    lung volume, an increased EMGdi signal intensity for the above    mentioned given inspiratory lung volume will increase the index but    will express a reduction in the neuro-ventilatory efficiency, and a    decreased EMGdi signal intensity for that given inspiratory lung    volume will reduce the index but will express an improvement of the    neuro-ventilatory efficiency.

In the following description, an example using the EMGdi signalintensity for a fixed, given inspiratory lung volume will be given.However, it is within the scope of the present invention to use the lungvolume for a fixed, given diaphragm EMGdi signal intensity.

Referring now to FIG. 10 a preferred, practical embodiment is described.A neuro-ventilatory efficiency computation device 601 receives thesignal 508 of FIG. 5 as well as the given, fixed inspiratory lungvolume. Device 601 comprises a unit 602 for determining the intensity ofthe signal 508 for the given inspiratory lung volume. Although it is notillustrated, it is within the scope of the present invention tocalculate, in unit 602, the peak, mean, median or any other intensitymeasure of signal 508 for the given inspiratory lung volume. If theintensity of signal 508 for the given inspiratory lung volume hasincreased at least by a given percentage (step 603), i.e. theneuro-ventilatory efficiency index has increased at least by said givenpercentage, the pressure, flow, or volume assist unit 604 is controlledby a unit 606 in view of increasing the magnitude of the pressure assistto the patient by a preset increment until the intensity of the signal608 for the given inspiratory lung volume is restored to apredetermined, preset value.

Still referring to FIG. 10, if the intensity for the given inspiratorylung volume has decreased at least by a given percentage (step 607),i.e. the neuro-ventilatory efficiency index has decreased at least bysaid given percentage, the pressure assist unit 604 is controlled by theunit 608 in view of decreasing the magnitude of the pressure assist by apreset increment until the intensity of the signal 508 for the giveninspiratory lung volume is restored to the predetermined, preset value.Although it is not illustrated, it is within the scope of the presentinvention to calculate, in unit 602, the peak, mean, median or any otherintensity measure of, signal 508 for the given inspiratory lung volume,instead of the intensity of this signal. Also, the signals at theoutputs of the units 606 and 608 can be used to generate an alarm or tomanually adjust the pressure, flow or volume assist to the patient.

The response time is adjustable. The time base used to calculate trendsin the EMG intensity for a given volume or vice versa and used for thecorrections is relatively slow (minutes) and the levels of appliedsupport can be limited within a safe range. Again, an alarm can begenerated or the pressure assist can be manually or automaticallyadjusted.

The pressure, flow, or volume assist unit 604 can be any device whichcan be controlled to generate any airway pressure of adjustablemagnitude, for example any source of compressed gas, or a flow or volumepump. Of course, airway 605 refers to or, to the least, includes thepatient's respiratory airway.

In this manner, the pressure assist unit 604 provides a pressure, flow,or volume assist that is adjusted in proportion to changed inneuro-ventilatory efficiency which is the EMGdi signal intensity at agiven lung volume or vice versa. The pressure, flow, or volume assistunit continuoulsy operates to maintain a tracheal pressure, flow orvolume that is adjusted in proportion to changes in neuro-ventilatoryefficiency which is the EMGdi signal intensity at a given lung volume orvice versa.

Pre-Inspiratory Breathing Effort

A common problem with mechanically ventilated patients is that thepatients' inspiratory effort will not immediately cause an inspiratoryairflow so called “intrinsic PEEP” or “auto PEEP” which leads to adecrease in the neuro-ventilatory efficiency. The effect of “intrinsicPEEP” can be counteracted by the application of an “extrinsic PEEP”.However, there are no easy applicable techniques to determine when theapplied level of “extrinsic PEEP” is adequate. The level ofpre-inspiratory effort obtained through the EMGdi signal intensity(common noise level subtracted) during for example a 100 milliseconds(ms) period immediately preceding the onset of inspiratory flow can beused to indicate the presence of “intrinsic PEEP”, and the level ofapplied “extrinsic PEEP” can be adjusted such that the level ofpre-inspiratory effort is suppressed i.e the EMGdi signal intensity(common noise level subtracted) during the above mentioned 100 ms periodbefore onset of inspiratory flow is close to zero. A feedback loop canthen be used to maintain the level of pre-inspiratory effort suppressedby adjusting as explained above the level of “extrinsic PEEP”.

Just a word to mention that the above mentioned period of 100 ms can bereplaced by a longer or shorter time period immediately preceding theonset of inspiratory flow or by the neuro-ventilatory delay 800 (FIG. 12b), i.e. the time period between the onset of EMG 801 (FIG. 12 b) andthe onset of inspiratory flow 802 (FIG. 12 a).

FIG. 11 of the appended drawings illustrates a preferred, practicalembodiment 700.

In the embodiment 700, an integrator 713 is responsive to the RMS EMGsignal 508 to continuously calculate the EMG intensity for the abovementioned 100 ms period or neuro-ventilatory delay 800.

Embodiment 700 also comprises an inspiratory flow detector 702responsive to the patient's inspiratory flow 703 measured, as indicatedin the foregoing description, through any commercially available system,to produce an output signal 705 representative of EMG activity.

The embodiment 700 of FIG. 11 also comprises a neuro-ventilatory delaycalculator 704 responsive to (a) the detection of a RMS EMG signalintensity higher than the common noise level (5%), and (b) the detectionof the onset of inspiratory flow by the detector 702 to calculate theneuro-ventilatory delay 800 (FIG. 12 b).

A detector 714 is responsive to the EMG intensity calculated by theintegrator 713 to detect the level of EMG intensity 803 (FIG. 12 b) atthe onset of inspiratory flow 802 (FIG. 12 a) to trigger an alarm 716when the level of the EMG intensity 803 at the onset of inspiratory flow802 is higher than a given limit (detector 715). Upon triggering of thealarm 716, the level of applied “extrinsic PEEP” is either automaticallyor manually increased (device 708).

The detector 714 is responsive to the EMG intensity calculated by theintegrator 713 to detect the level of EMG intensity 803 (FIG. 12 b) atthe onset of inspiratory flow 802 (FIG. 12 a) to trigger an alarm 720when the level of the EMG intensity 803 at the onset of inspiratory flow802 is lower than a given limit (detector 719). Upon triggering of thealarm 720, the level of applied “extrinsic PEEP” is either automaticallyor manually decreased (device 711).

It should be mentioned that feedback from the neuro-ventilatory delay orpre-inspiratory EMG activity can also be used to adjust the sensitivityof the ventilators trigger functions.

Again, the time base used for these corrections is preferably relativelyslow (minutes) and the levels of “extrinsic PEEP” can be limited withina safe range.

The pressure assist unit 604 can be any device which can be controlledto generate any airway flow and/or pressure of adjustable magnitude, forexample any source of compressed gas, or a flow or volume pump.

In this manner, the delay from the beginning of the mechanicallyventilated patients' inspiratory effort to the onset of the inspiratoryassist will be minimized.

Although the present invention has been described hereinabove withreference to preferred embodiments thereof, these embodiments can bemodified at will, within the scope of the appended claims, withoutdeparting from the spirit and nature of the subject invention.

1. A neuro-ventilatory efficiency computation method formonitoring/controlling a level of ventilatory assist to a patientcomprising: receiving an EMG signal intensity representative ofinspiratory effort of the patient receiving a lung volume valuerepresentative of a lung volume of the patient determining from thereceived EMG signal intensity and lung volume value at least one of thetwo following relations: an EMG signal intensity for a given lung volumevalue, the received lung volume value then being said given lung volumevalue; and a lung volume value for a given EMG signal intensity, thereceived EMG signal intensity then being said given EMG signalintensity; and increasing or decreasing the ventilatory assist leveldepending on whether said at least one relation has increased ordecreased by at least a given percentage.
 2. A neuro-ventilatoryefficiency computation method as defined in claim 1, wherein: increasingor decreasing the ventilatory assist level comprises increasing theventilatory assist level by a preset increment when said at least onerelation has increased by at least said given percentage.
 3. Aneuro-ventilatory efficiency computation method as defined in claim 1,wherein: increasing or decreasing the ventilatory assist level comprisesincreasing the ventilatory assist level by a preset increment when saidat least one relation has increased by at least said given percentageuntil the EMG signal intensity for the given lung volume value isrestored to a predetermined, preset value.
 4. A neuro-ventilatoryefficiency computation method as defined in claim 1, wherein: increasingor decreasing the ventilatory assist level comprises decreasing theventilatory assist level by a preset decrement when said at least onerelation has decreased by at least said given percentage.
 5. Aneuro-ventilatory efficiency computation method as defined in claim 1,wherein: increasing or decreasing the ventilatory assist level comprisesdecreasing the ventilatory assist level by a preset decrement when saidat least one relation has decreased by at least said given percentageuntil the EMG signal intensity for the given lung volume value isrestored to a predetermined, preset value.
 6. A neuro-ventilatoryefficiency computation method as defined in claim 1, wherein increasingor decreasing the ventilatory assist level comprises: increasing theventilatory assist level by a preset increment when said at least onerelation has increased by at least said given percentage; and decreasingthe ventilatory assist level by a preset decrement when said at leastone relation has decreased by at least said given percentage.
 7. Aneuro-ventilatory efficiency computation method as defined in claim 1,further comprising: generating an alarm when said at least one relationhas increased or decreased by the given percentage.
 8. Aneuro-ventilatory efficiency computation method as defined in claim 1,further comprising: manually adjusting the ventilatory assist level inresponse to a signal from the operation of increasing or decreasing theventilatory assist level.
 9. A neuro-ventilatory efficiency computationmethod as defined in claim 1, wherein: determining at least one relationcomprises calculating one of the following values of the EMG signalintensity or lung volume value: a mean of the EMG signal intensity orlung volume value, a median the EMG signal intensity or lung volumevalue, and a peak the EMG signal intensity or lung volume value.
 10. Aneuro-ventilatory efficiency computation method as defined in claim 1,wherein the EMG signal intensity is a patient's diaphragm EMG signalintensity.
 11. A neuro-ventilatory efficiency computation method asdefined in claim 1, comprising: calculating a trend in the EMG signalintensity for a given lung volume value using an adjustable time base.12. A neuro-ventilatory efficiency computation method as defined inclaim 1, comprising: calculating a trend in the lung volume value for agiven EMG signal intensity using an adjustable time base.
 13. Aneuro-ventilatory efficiency computation method as defined in claim 1,comprising: limiting a range of the ventilatory assist level within asafe range.
 14. A neuro-ventilatory efficiency computation device formonitoring/controlling a level of ventilatory assist to a patientcomprising: a first input for receiving an EMG signal intensityrepresentative of inspiratory effort of the patient; a second input forreceiving a lung volume value representative of a lung volume of thepatient; connected to the first and second inputs a calculator of atleast one of the two following relations: an EMG signal intensity for agiven lung volume value, the lung volume value received on the secondinput then being said given lung volume value; and a lung volume valuefor a given EMG signal intensity, the EMG signal intensity received onthe first input then being said given EMG signal intensity; and acontroller connected to the first unit, the controller increasing ordecreasing the ventilatory assist level depending on whether said atleast one relation has increased or decreased by at least a givenpercentage.
 15. A neuro-ventilatory efficiency computation device asdefined in claim 14, wherein: the controller increases the ventilatoryassist level by a preset increment when said at least one relation hasincreased by at least said given percentage.
 16. A neuro-ventilatoryefficiency computation device as defined in claim 14, wherein: thecontroller increases the ventilatory assist level by a preset incrementwhen said at least one relation has increased by at least said givenpercentage until the EMG signal intensity for the given lung volumevalue is restored to a predetermined, preset value.
 17. Aneuro-ventilatory efficiency computation device as defined in claim 14,wherein: the controller decreases the ventilatory assist level by apreset decrement when said at least one relation has decreased by atleast said given percentage.
 18. A neuro-ventilatory efficiencycomputation device as defined in claim 14, wherein: the controllerdecreases the ventilatory assist level by a preset decrement when saidat least one relation has decreased by at least said given percentageuntil the EMG signal intensity for the given lung volume value isrestored to a predetermined, preset value.
 19. A neuro-ventilatoryefficiency computation device as defined in claim 14, wherein thecontroller: increases the ventilatory assist level by a preset incrementwhen said at least one relation has decreased by at least said givenpercentage; and decreases the ventilatory assist level by a presetdecrement when said at least one relation has decreased by at least saidgiven percentage.
 20. A neuro-ventilatory efficiency computation deviceas defined in claim 14, further comprising: an alarm generated when saidat least one relation has increased or decreased by the givenpercentage.
 21. A neuro-ventilatory efficiency computation device asdefined in claim 14, further comprising: means for manually adjustingthe ventilatory assist level.
 22. A neuro-ventilatory efficiencycomputation device as defined in claim 14, wherein: the calculatordetermines one of the following values of the EMG signal intensity orlung volume value: a mean of the EMG signal intensity or lung volumevalue, a median the EMG signal intensity or lung volume value, and apeak the EMG signal intensity or lung volume value.
 23. Aneuro-ventilatory efficiency computation device as defined in claim 14,wherein the EMG signal intensity is a patient's diaphragm EMG signalintensity.
 24. A neuro-ventilatory efficiency computation device asdefined in claim 14, comprising: means for calculating a trend in theEMG signal intensity for a given lung volume value using an adjustabletime base.
 25. A neuro-ventilatory efficiency computation device asdefined in claim 14, comprising: means for calculating a trend in thelung volume value for a given EMG signal intensity using an adjustabletime base.
 26. A neuro-ventilatory efficiency computation device asdefined in claim 14, comprising: means for limiting a range of theventilatory assist level within a safe range.